In the field of nuclear medical imaging technology, which includes PET imaging detector 10, as illustrated in FIG. 1, an array of radiation sensors, such as plurality of scintillators 13 and associated photosensors 11n, such as photomultiplier tubes (PMTs), avalanche photodiodes (APDs), or silicon photomultipliers (SiPMs) are usually arranged in a circle of a detector ring 13. Such a detector ring 13 surrounds a subject to be scanned. To conduct a so-called PET scan, a short-lived radioisotope, which decays by emitting a positron, is injected usually into the blood circulation of a living subject. After the metabolically active molecule becomes concentrated in tissues of interest, the research subject or patient is placed in the imaging scanner. The most commonly-used metabolically active molecule for this purpose is fluorodeoxyglucose (FDG), a sugar, which has a half life of 110 minutes. Note that other radiation sensors, such as solid state detectors could be used in the place of scintillators and photosensors.
As the radioisotope undergoes positron emission decay, it emits a positron, the antimatter counterpart of an electron. After traveling up to a few millimeters, the positron encounters and annihilates with an electron, producing a pair of gamma photons moving in almost opposite directions. These are detected when they reach one of a plurality of scintillation crystals in the scanning device, creating a burst of light detected by an array of photosensors. These bursts of light from a scintillator, such as Lutetium Orthosilicate (LSO) and Bismuth Germanate (BGO), have an intrinsic shape with a fast rising edge followed by a slow falling edge. The signals can be estimated as a function of:
                                          V            o                    ⁡                      (            t            )                          ≈                              A            1                    ·                      m            0                    ·                      (                                                            1                                      τ                    1                                                  ·                                  ⅇ                                                            -                      t                                        /                                          τ                      1                                                                                  -                                                1                                      τ                    0                                                  ·                                  ⅇ                                                            -                      t                                        /                                          τ                      0                                                                                            )                                              (        1        )            where τ0 is the characteristic scintillator decay time constant; and τ1 is mainly determined by the characteristics of the photosensor (such as PMT, APD, or SiPM), the open-loop gain of the first amplifier in the front-end electronics, and the input capacitance. When τ0>>τ1 (which are the cases for LSO and BGO crystals), τ1 dominates the rising edge of the pulse, and τ0 dominates the falling edge.
The Laplace transfer-function of Eq. 1 is:
                              H          ⁡                      (            s            )                          =                                            A              1                                      τ              1                                ·                      s                                          (                                  s                  +                                      1                                          τ                      0                                                                      )                            ·                              (                                  s                  +                                      1                                          τ                      1                                                                      )                                                                        (        2        )            
As can be seen in Eq. 2, the falling edge of the scintillation signal is a first-order exponential decay function, so the shape of the signal is always unipolar—it is either positive or negative depending on the polarity chosen for the sensor analog electronics.
Generally a plurality of photosensors 11 can be arranged in a matrix and assigned to detect the light of a single scintillator as shown in the enlarged section 16 in FIG. 1. The scintillator detector 13 can comprise a single scintillation crystal or can be, as shown, a matrix of scintillator crystals that is coupled to the photosensors 111 . . . 11n via a light guide 17. To be able to increase the resolution of the system without the high costs of 1:1 coupling, the number of photosensors 11 per block is generally significantly lower than the number of scintillation crystals 13. For example, a block detector may have a plurality of photosensors 11 with, for example, 4, 9 or 16 photosensors 11 arranged in a 2×2, 3×3, or 4×4 matrix behind an array of scintillation crystals 13. Other arrangements with more or fewer photosensors 11 are possible. Thus, scintillation event localization can be determined or interpolated by such a scintillation block detector by processing the associated photosensor signals. This can be done by analog filtering, integration, and multiplication of weighted combinations of the photosensor signals or by using digital algorithms that use discrete time sample points of signals obtained directly from the photosensors 11. The PET technique depends on scintillation event detection of the pair of gamma photons.
FIG. 1 illustrates a block diagram of the typical architecture of detectors and associated analog-to-digital-converters in a conventional system. Each matrix of photosensors 11 produces a plurality of signals that can be processed to generate an image from a plurality of scintillation events that are detected by a photosensor 11. To determine the location of a detected annihilation, the system needs to accurately measure the timing and energy of both detected photons. Consequently a high amount of data has to be produced by the respective measurement circuits.
For example, as shown on the right side of FIG. 1, each scintillator has an associated matrix of detectors, such as photosensors 111 . . . 11n, in this example are PMTs for illustration. Each signal of each PMT 111 . . . 11n is first amplified by, for example, associated preamplifiers/buffers 121 . . . 12n. The output signal of preamplifier/buffers 121 . . . 12n can then be converted concurrently into discrete-time digital signals by associated analog-to-digital converters (ADC) 141 . . . 14n. A sampling clock for each ADC be can provided at terminal 15. In this example, this digital processing architecture uses n independent ADC signals with peripheral circuitry to concurrently sample each of n photosensor signals per block. This can increase the costs of a detector unit.
FIG. 2 illustrates a detector scintillation block comprising an 8×8 array of scintillation crystals 131-13n; for example, each crystal can be 4 mm×4 mm×20 mm. Photosensors 11 can be included behind the scintillation crystals to detect light emitted due to scintillation events. Exemplary photosensor arrangements of photosensors are illustrated in FIGS. 3 and 4. FIG. 3 illustrates photosensors 11 in a 3×3 array and FIG. 4 illustrates photosensors 11 in a 4×4 array. Either of these arrangements, and others, could be used to detect light emitted from the 8×8 array of scintillation crystals illustrated in FIG. 2.